Field of the Invention
The present invention concerns a method and a control device to control a magnetic resonance system to execute a pulse sequence. The invention also concerns a magnetic resonance tomography system, also called a magnetic resonance system in the following.
Description of the Prior Art
Magnetic resonance tomography—also called nuclear magnetic resonance tomography—is a widespread technique to acquire images of the inside of the body of a living examination subject.
Atomic nuclei, for example of hydrogen atoms, exhibit a property known as a spin, which is a quantum mechanical property of atomic particles. The spin has the effect that the atomic particles are magnetic dipoles, meaning that an atomic nuclei with spin are magnetic dipoles. These spins initially act in any direction. They can be considered as a vector. Atoms with spin are present in a body to be examined, for example a human body.
In a magnetic resonance tomography system, the body to be examined is typically exposed to a relatively high basic field magnet field B0 (for example of 1, 5, 3 or 7 Tesla) with the use of a basic field magnet system. The force effect of the static magnetic field B0 generates a preferred direction of the spins parallel and antiparallel to the field lines. An excess always forms in one direction, which leads to a macroscopic magnetization of the spin ensemble.
A radio-frequency magnetic field B1 is superimposed on the static magnetic field B0. This radio-frequency magnetic field (which normally is generated by radio-frequency excitation pulses) brings the spins out of the steady state generated by the B0 field when the radio-frequency excitation signals are in resonance with the precession frequency of the spins. The precession frequency is also called Larmor frequency. It is dependent on the strength of the external magnetic field. By means of the radio-frequency excitation signals, the nuclear spins of the atoms excited to resonance by this radio-frequency field are flipped by a defined flip angle relative to the magnetic field lines of the basic magnetic field.
The emission of the radio-frequency signals for nuclear resonance magnetization most often takes place by means of what is known as a “whole-body coil” or “body coil”. A typical design of a whole-body coil is a cage antenna (birdcage antenna) which comprises multiple transmission rods that—running parallel to the longitudinal axis—are arranged around a patient space of the tomograph in which a patient is located in the examination. The antenna rods are respectively capacitively connected with one another in an annular form on their front sides. However, currently local coils close to the body are being used more often for the emission of MR excitation signals. The reception of magnetic resonance signals normally takes place with the local coils, but in some cases also alternatively or additionally with the body coil.
A gradient is applied along a gradient direction via gradient coils. The magnetic field B0 thereby increases linearly. The precession of the nuclear spins along the gradient direction is accordingly different; the spins spin slower here, faster there. They therefore show resonance at different frequencies. A spatially selective excitation of the nuclear spins is possible via the superimposed gradient field.
The exciting radio-frequency signal or the exciting radio-frequency pulse receives a defined bandwidth of neighboring frequencies around a center frequency. In this way a desired region along the gradient direction can be excited.
In nearly all molecules, multiple hydrogen atoms are bound at various positions. Various positions mean different chemical (and therefore most often also different magnetic) environments. The local magnetic field is hereby reduced or, respectively, increased; the resonance frequencies of the bound protons are somewhat lower or higher than the typical Larmor frequency.
The nuclear spins in the body tissue thus do not have a uniform precession frequency in the magnetic field, but rather differ according to their chemical environment for different tissue types. This is typically designated as a chemical shift. Fat has multiple peaks in the spectrum, but one is strongly pronounced and delivers a high signal for imaging. The chemical shift between the primary peak of the adipose tissue and water is approximately 3.5 ppm, for example.
After an excitation, the nuclear spins flip back again into their initial state that is enforced by the basic magnetic field. This is what is known as the relaxation of the nuclear spins. It is differentiated in longitudinal relaxation and transversal relaxation. The longitudinal relaxation describes the re-establishment of the magnetization along the magnetic field lines of the basic magnetic field B0. The transversal relaxation describes the disappearance of the magnetization caused by the radio-frequency field B1 transversal to the magnetic field lines of the basic magnetic field B0. Different tissues have different relaxation times.
In the precession, radio-frequency signals (what are known as magnetic resonance signals) are radiated that are received and processed further by means of suitable reception antennas. The desired image data are reconstructed from the raw data acquired in such a manner. The reception antennas can either be the same antennas with which the radio-frequency excitation pulses are also radiated or separate reception antennas.
As used herein, the term “fat signal” means the signal that a nuclear spin that is situated in adipose tissue emits upon relaxation. The term “water signal” means the signal that a nuclear spin that is located in an aqueous region emits upon relaxation.
The signals emitted in the precession and received by the reception antennas must be capable of spatial association in order to enable an imaging. For this, a spatial coding is implemented in the acquisition of the signals via coding gradients.
Given 2D magnetic resonance pulse sequences, a spatial coding takes place in two directions or dimensions. Therefore, image information or raw data are respectively read out for a very thin slice. The slice is selected beforehand. Given 3D magnetic resonance pulse sequences, a spatial coding takes place in three directions or dimensions. Therefore, image information or raw data are respectively read out for an entire volume, what is known as a “slab”.
The raw data are written into a matrix in an electronic memory known as k-space. K-space is a space or a spatial frequency domain that is Fourier-transformed into a positional space that includes the subject magnetization. The axes of k-space designate what are known as spatial frequencies. K-space has a unit that is inverse to the distance, for example 1/cm. In 3D tomography, k-space is also three-dimensional.
Static magnetic field differences contribute to a diversification of the spins upon relaxation. With spin echo sequences, this diversification is canceled via a refocusing pulse or, respectively, via a series of refocusing pulses. If multiple refocusing pulses (normally 180° pulses) follow in series, multiple spin echoes arise, generated by a multi-echo sequence. The registration in k-space depends on the desired contrast, among other things. Often the earlier echoes—i.e. the echoes with a smaller position number—are initially registered in central k-space.
SPACE (Sampling Perfection with Application optimized Contrast using different flip angle Evolutions) is an example of a three-dimensional turbo spin echo sequence method—more precisely a single slab 3D turbo spin echo method—that can have very long echo trains. For example, a long echo time includes between forty and multiple hundreds of echoes; thousands of echoes are also possible, for instance. For a “provided signal development” (prescribed signal evolution), the flip angle of the refocusing pulses in an echo train is adapted to the properties (T1 and T2) of the different tissue types. A variable flip angle curve (flip angle evolution) is obtained. A desired signal strength is generated for different types of tissue. For example, a desired contrast can therefore be generated.
The magnetic resonance images of the examination subject are ultimately created on the basis of the received magnetic resonance signals. Each image point in the magnetic resonance image is thereby associated with a small physical volume—what is known as a “voxel”—and each brightness or intensity value of the image points is linked with the signal amplitude of the magnetic resonance signal that is received from this voxel. The connection between a resonant, radiated RF pulse with field strength B1 and the flip angle α that is therefore achieved is thereby provided by the equation
                              α          =                                    ∫                              t                =                0                            τ                        ⁢                          γ              ·                                                B                  1                                ⁡                                  (                  t                  )                                            ·                                                          ⁢                              ⅆ                t                                                    ,                            (        1        )            wherein γ is the gyromagnetic ratio—which for most magnetic resonance examinations can be viewed as a fixed material constant—and τ is the effective duration of the radio-frequency pulse.
In many cases the very bright fat signal, which in many cases outshines the water signal of primary interest, is problematic in the diagnosis of possible pathologies.
Therefore, possibilities have already been proposed to suppress the fat signal. For example, before the actual measurement a frequency-selective pulse is emitted at the precession frequency of the protons situated in adipose tissue so that their spins are saturated and do not contribute to the signal in the subsequent image acquisition.
A further possibility is offered by what is known as the Dixon sequence, in which multiple echoes are acquired at different echo times (time after an excitation or, respectively, refocusing pulse). The image data of different materials (thus for example fat and water) are shifted in their phase. It is a post-processing measure, meaning that the acquired raw data are retroactively processed. Such a Dixon sequence is, for example, described in H. Yu et al.: “Implementation and Noise Analysis of Chemical Shift Correction for Fast Spin Echo Dixon Imaging” in Proc. Intl. Soc. Mag. Reson. Med. 11 (2004), 2686, wherein a Dixon sequence is used in the method described there in order to achieve a correction of the chemical shift within an image plane via multiplication of the k-space lines with a defined phase term.
However, the methods that have previously been known have disadvantages in regions with strong B1 inhomogeneities: conventional fat saturation methods are based on the fact that the predetermined flip angle is achieved as exactly as possible. At high basic field strengths (B0≧3T), this is often not the case. Since the proposed method foregoes pre-pulses and inversion pulses, this is less susceptible to B1 inhomogeneities. Known methods are also limited in the image resolution that can be achieved: under the circumstances, the desired resolution cannot be achieved in Dixon TSE due to timing problems of the individual echoes in the echo train. Confusion can also occur in the separation of fat and water in the DIXON method, such that the fat image incorrectly shows a water image and vice versa.